Electric field probe

ABSTRACT

An electric field probe array including several dipole probes, each including a diode connected between two conductors with the diodes being collectively connected to at least two probe terminals. Also disclosed is a system for heating a target with electromagnetic radiation that includes a central control unit, a source coupled to the control unit with the source generating the electromagnetic energy at a frequency and power in response to the control unit, a radiation apparatus coupled to the source for transferring the electromagnetic energy from the source to the target and input circuitry connected to the control unit for indicating the status of the target whereby the operation of the system is controlled by the control unit as a function of the target&#39;s status. The input circuitry includes several electric field probe arrays where each electric field probe array includes several individual diode probes that each include a diode connected between two conductors with the diodes collectively connected to electric field probe array terminals which are connected to the input circuitry.

RELATED APPLICATIONS

The present application is a continuation-in-part of U.S. applicationSer. No. 405,947, filed Aug. 6, 1982 which is a continuation-in-part ofU.S. application Ser. No. 136,506, filed Apr. 2, 1980, now U.S. Pat. No.4,462,412.

BACKGROUND

1. Field of the Invention

The present invention relates generally to systems for irradiatingtargets with electromagnetic radiation, and more specifically toannular-type applicators and associated systems for controllingapplication of radiation to biological tissue.

2. Description of Prior Art

Several types of therapeutic treatments for cancer in humans are incurrent, common use. These treatments include surgery, X-rays, radiationfrom radioactive sources and chemotherapy, these treatments being oftencombined in various ways to enhance treatment effectiveness.

Although such conventional treatment techniques have been successful intreating cancer in many patients and in prolonging the lives of manyother patients, they are frequently ineffective against many types ofcancer and often have severe adverse side effects at the necessarytreatment levels. Protracted treatment of cancer patients by x-rays orchemotherapy, as an illustration, tends to eventually destroy or inhibitthe patients' natural immunological systems to an extent that manypatients eventually succumb to common infectious diseases, such asinfluenza or pneumonia, which otherwise probably would not be fatal.Also, many patients having advanced stages of cancer or complicationsmay become too weak to withstand the trauma of surgical or other cancertreatments; hence, the treatments cannot be undertaken or must bediscontinued.

Due both to the prevalence and the typically severe consequences ofhuman cancer, as well as frequent ineffectiveness of current treatmentssuch as those mentioned above, medical researchers are continuallyexperimenting in an attempt to discover and develop improved oralternative cancer treatment methods with their associated treatmentapparatus.

Hyperthermia, the generation of artificially elevated body temperatures,has recently been given serious scientific consideration as analternative cancer treatment. Much research has been conducted into theeffectiveness of hyperthermia alone or in combination with othertreatment methods. This research is important in that hyperthermiatechniques appear to have the potential for being extremely effective inthe treatment of many or most types of human cancers, without the oftenseverely adverse side effects associated with current cancer treatments.

Researchers into hyperthermia treatment of cancer have commonly reportedthat many types of malignant growths in humans can be thermallydestroyed, usually with no serious adverse side effects, by heating themalignancies to temperatures slightly below that injurious to mostnormal, healthy cells. Furthermore, many types of malignant cell masseshave reportedly been found to have substantially poorer than normal heattransfer or dissipation characteristics, presumably due to poorervascularity and reduced blood flow characteristics. Consequently, thesetypes of growths appear capable of preferential hyperthermia treatment.Poorly vascular malignant growths can reportedly be heated totemperatures several degrees higher than that in which the immediatelysurrounding healthy tissue reaches. This promises to enable hyperthermictreatment of those types of malignant growths which are no morethermally sensitive than normal tissue without destruction of normalcells, and additionally to enable higher temperature, shorterhyperthermia treatment times of more thermally sensitive types ofmalignancies which exhibit poor vascularity, usually an advantage forimportant medical reasons.

In this regard, researchers have commonly reported that as a consequenceof these thermal characteristics of most malignant growths and thethermal sensitivity of normal body cells, hyperthermia temperatures fortreatment of human cancer should be carefully limited within arelatively narrow effective and safe temperature range. Below athreshold temperature of about 41.5° C. (106.57° F.), significantthermal destruction of most malignant growths of cells has normally notbeen found to occur.

At slightly higher hyperthermic temperatures, above the approximaterange of 43° C. to 45° C. (109.4° F. to 113° F.), thermal damage to mosttypes of normal cells is routinely observed; thus, great care must betaken not to exceed these temperatures in healthy tissue. Exposureduration at any elevated temperature is, of course, an important factorin establishing extent of thermal damage to healthy tissue. However, iflarge or critical regions of the human body are heated into, or abovethe, 43° C. to 45° C. range, for even relatively short times, seriouspermanent injury or death may be expected to result.

Historically, alternating electric currents, at frequencies above about10 KHz, were found late in the last century to penetrate and causeheating in biological tissue. As a result, high frequency electriccurrents, usually in the megahertz frequency range, have since beenwidely used for therapeutic treatment of such common bodily disorders asinfected tissue and muscle injuries. Early in this century, the name"diathermia" was given to this EMR tissue heating technique, and severaldiscrete EMR frequencies in the megahertz range have subsequently beenallocated specifically for diathermy use in this country by the FederalCommerce Commission (FCC).

A number of even more current discussions relating to EMR hyperthermiatreatment of cancer may, for example, be found in a compilation ofarticles on the subject published in the book "Cancer Therapy byHyperthermia and Radiation", edited by Christian Streffer et al andpublished by Urban and Schwarzenberg; Baltimore, Munich 1978.

In spite of there having been reported encouraging and often apparentlysuccessful medical results obtained by using EMR induced hyperthermia totreat malignant growths in humans, the treatments have normally been ofan experimental nature, typically being used on cancer patientsotherwise considered incurable or terminal, since serious problemsrelating to hyperthermic damage to healthy tissue have commonly beenencountered. As with conventional surface heating, these healthy tissuedamage problems are particularly associated with thermally destroyingmalignant growths deeply located in, or close to, thermally sensitivetissue.

This unintended EMR thermal damage of healthy tissue can typically beattributed to design and use of existing EMR irradiating apparatus,rather than to any basic deficiency in the concept of EMR hyperthermiatreatment. EMR apparatus used, for example, often radiate excessiveand/or improperly controlled EMR heating fields. A further disadvantageis that the specific diathermy allocated frequencies which areordinarily used are typically non-optimum radiating frequencies for deeppenetration. In addition, existing EMR hyperthermia apparatus andtechniques tend to increase incidence and severity of thermal "hotspotting" in healthy tissue, as may be caused by uncontrolledconstructive interference of applied energy waves, either bycharacteristic reflections at interfaces between different types oftissue, or by simultaneous use of more than one EMR applicator.

To overcome these and other problems associated with heretoforeavailable EMR hyperthermia apparatus used to medical research or othermedical purposes, applicant has disclosed improved EMR hyperthermiaapparatus in U.S. patent applications, Ser. Nos. 002,584, now U.S. Pat.No. 4,271,848, and 048,515, now U.S. Pat. No. 4,341,227, filed on Jan.11, 1979 and June 14, 1979, respectively. In these two patentapplications, parallel plate and waveguide-type EMR applicators,together with associated EMR systems, were described and claimed, theapplicators being particularly adapted for irradiating biological tissueor tissue simulating matter from outside the tissue. Emphasis was placedon broad band EMR capabilities, enabling, for example, researchdefinition of important parameters associated with hyperthermictreatment of malignancies in humans. Also described in such patentapplications was simultaneous operation of two (or more) applicatorsarranged to improve deep tissue heating characteristics.

In applicant's subsequent U.S. patent application, Ser. No. 050,050,filed on June 19, 1979, now abandoned, needle-type invasive EMRapplicators, for enabling EMR hyperthermia in sub-surface tissueregions, were described. By surrounding, with a phased array of theseinvasive applicators, a localized tissue region, such as a regioncontaining a malignant growth, substantially uniform heating of thesurrounded region by constructive interferences of the synchronous EMRfiled was described.

However, there still exists an important need for EMR hyperthermiaapparatus capable of causing uniform deep EMR heating of thick tissuemasses, such as trunk and thigh portions of an adult human body, inwhich large or widely dispersed malignant growths may be found. Forthese and similar regions of the body, an encircling annular EMRApplicator apparatus, which may comprise an array of smallerapplicators, is odinarily preferred so that EMR energy is emittedinwardly from all around the enclosed body region to be EMR heated.

To this end, large annular magnetic coils have been used to radiate amagnetic field into a body region disposed through the coil. Althoughsuch radiated magnetic field are known to penetrate deeply in humantissue, uniform heating across the encircled tissue region is normallynot possible. This has been reported by several researchers. This isbecause the induced currents couple much stronger in the longer currentpaths along the outer tissue regions than in the center. Thus, when suchannular magnetic coils are used to cause hyperthermia in biologicaltissue, near surface tissue regions can be expected to be heated muchmore than underlying central tissue regions.

Presently available EMR applicators have failed to permit deep heatingof a large target, such as a human torso, without severely damaging theoverlying tissue. This is due to the lossy nature of biological tissue,which absorbs a certain fraction of EMR passing therethrough, suchfraction being dependent on the tissue type and frequency of the EMR.Large amounts of the available power are absorbed near the surface,leaving relatively small amounts to be absorbed in the deeper regions.This causes excessive tissue heating at the surface, with littletherapeutic heating in the desired location, which is the central regionof the target.

Some heating improvement can be had in the central heating region byusing several applicators surrounding the target. Such an arraytypically energizes the individual applicators at different frequencies,or frequencies which vary slightly with time, in order to avoidundesired hot regions caused by constructive interference between thedifferent wavefronts. Such arrangements increase the power density inthe central region to approximately the sum of the power densities dueto the individual applicators without increasing appreciably the energyabsorbed in a unit area of the target's surface. However, the powerdensity in near-surface regions is usually still such as to be harmfulto biological targets before heating reaches therapeutic values in thedeeper regions.

When success in thermally treating deeply located malignant growths byheretofore available types of annular applicator apparatus has beenexperienced, the results appear to be more attributable to poorer heatdissipation properties of large malignant growths than to the desirableuniformity of heating. Heating uniformity, not heretofore available,would also enable effective thermal treatment, for example, of deeplylocated, widely dispersed, small groups of malignant cells in earlystages of growth before the associated reduced heat transfercharacteristics become significant. Ability to provide at leastsubstantially uniform heating of encircled tissue regions is also veryimportant, for example in areas of EMR hyperthermia research intohyperthermic effect on normal healthy tissue.

SUMMARY OF THE INVENTION

In accordance with the present invention an electric field probe arrayis disclosed that includes a plurality of dipole probes where eachdipole probe includes a diode connected between two conductors which arecollectively connected to at least two terminals.

In a preferred embodiment of the present invention, an electric fieldprobe array is disclosed that includes several of the diode probescollectively connected to at least two plurality conductive leads of aresistive value per length preferably greater than (typically ten timesgreater than) the resistance per length of a single resistive dipoleprobe. These partially conductive leads are connected to the probeoutput monitor terminals. In this embodiment the conductors connected tothe diodes are axially aligned and the conductors are of a conductivematerial that has a characteristic resistance of greater than 500 ohmsper centimeter to minimize lead heating in an electromagnetic field. Theelectric field probe is for use in systems that heat a target masscontaining pertubating nonhomogeneous structures. In this application,the length of the dipole array is larger than any single pertubatingnonhomogeneous structures within the target and each dipole is of a sizethat is smaller than one third of a wavelength of any dominant materialwithin this target to avoid dipole resonant effects.

In a further embodiment, a system is disclosed for heating a target withelectromagnetic radiation that includes a central control unit and asource connected to the control unit for generating electromagneticenergy at a frequency and power in response to the control unit. Aradiation apparatus is connected to the source for transferring thiselectromagnetic energy from the source to the target. Input circuitry isconnected to the control unit for indicating the electromagneticelectric field amplitude at the target tissue whereby the operation ofthe system is controlled manually or by the control unit as a functionof the E-fields surrounding the target tissue. This input circuitryincludes several electric field probe arrays where each electric fieldprobe array individually includes several dipole probes. Forelectromagnetic sources such as the annular array which has steerableheat patterns these probes can direct the beam steering under CPUcontrol. Each dipole probe together includes a diode connected betweentwo conductors with the diodes collectively connected to terminalsconnecting to the input circuitry. In the preferred embodiment, theelectric field probes may include catheters for inserting the electricfield probe arrays into or near the target. Also in this embodiment, thedipole conductors are of a length below one third the wavelength offrequency of any individual structure contained within the target. Thesource frequency typically exceeds 40 megahertz. The diodes of theelectric field probe operate within the linear region within theircharacteristic curves and preferably are zero bias Schotky diodes. Ifoperated in the non-linear range the CPU can correctly display theelectric field amplitudes. The diodes are collectively connectedtogether with a resistive lead that has a resistive characteristic of atleast 600 ohms per centimeter of length to assure heating of the lead inelectromagnetic fields are minimal. In this embodiment the probe arrayis sealed with a dielectric insulator.

BRIEF DESCRIPTION OF THE DRAWINGS

The novel features which characterize the present invention are definedby the appended claims. The foregoing and other objects and advantagesof the invention will hereinafter appear, and for purposes ofillustration, but not of limitation, a preferred embodiment is shown inthe accompanying drawings.

FIG. 1 is a schematic diagram of a system for creating hyperthermia in atarget;

FIG. 2 is a diagram showing the method of operation of a systemaccording to FIG. 1;

FIG. 3 is a diagram showing relative electric field amplitudes within ahomogeneous target specimen;

FIG. 4 is a diagram showing relative power density in a homogeneoustarget specimen;

FIG. 5 is a partially cut away perspective view of one preferredapplicator for use with the present invention;

FIG. 6 is a top diagrammatic view showing operation of the applicator ofFIG. 5;

FIG. 7 is a diagram of a dipole antenna for use with the presentinvention;

FIG. 8 is a perspective view of a folded dipole array for use with thesystem of the present invention;

FIG. 9 is a perspective view of a third preferred applicator for usewith the present invention;

FIG. 10 is an idealized diagrammatic view illustrating the operation ofmovable temperature probes for use with the system of the presentinvention; and

FIG. 11 is a cross section of an idealized torso used as a targetspecimen in conjunction with the system of the present invention,showing the location of electric field detectors.

FIG. 12 is a schematic diagram of the electric field probe array.

FIG. 13 is a block diagram of a system including an electric field probearray and a receiving apparatus.

DESCRIPTION OF THE PREFERRED EMBODIMENTS

FIG. 1 shows a block diagram of a system 10 for creating hyperthermia ina target specimen 12 by means of electromagnetic radiation (EMR). Acentral processor unit 14 controls the system 10, and is in interactivefeedback relationship with each of its elements. The central processorunit (CPU) 14 accepts a plurality of inputs describing the presentcondition of the target 12. These inputs can include, but are notnecessarily limited to, signals relating to the target's vital signs 16,internal temperature distribution 18, and external electric fielddistribution 20.

The vital signs input 16 is preferably for use with live targets 12.These signs can include, for example, mean arterial, diastolic, andsystolic blood pressure, heart rate, oral temperature and other signsgenerally used by medical personnel to determine the overall status of apatient.

The CPU 14 will also accept inputs 18 from a plurality of thermal probes22, which is diagrammatically shown in FIG. 1 as a single probe 22.These probes 22 give a real time indication of the internal temperatureswithin the target 12, and, when probes 22 of a type such as thosedescribed in connection with FIG. 10 are used, can give a either a twodimensional or three dimensional real time representation of thermaldistributions within the target 12. This thermal information can be usedto ensure that desired portions of the target 12 are heated to theproper temperature, and that no undersirable thermal gradients orconcentrations occur.

The CPU 14 can also accept inputs 20 from electric field detectors 24which indicate the real time electric field amplitude at selectedlocations. These detectors 24 are normally external of the target 12,and are preferably placed against or close to the surface. A pluralityof E-field detectors 24 can be used, preferably at least 8, and theseare represented diagrammatically by a single input 20 to the CPU 14.These E-field detectors 24 provide feed-back to the CPU 14 of therelative amplitude balance of the EMR incidental to the target surfacethereby allowing control of heat pattern by steering using relativephase, amplitude, and tissue coupling.

A control panel or console 26 is coupled to the CPU 14 and is used by anoperator to control a treatment and monitor its progress. The controlpanel 26 can be used to display any information obtained from the target12 as well as all indicators of system operation. Various memorydevices, represented by a single memory block 28, are coupled to the CPU14. The memory 28 stores the result of pretreatment calculations whichare used by the CPU 14 to control the progress of the treatment. Also,all pertinent operating data are stored in another part of the memory 28as generated in order to have a complete record of the treatment processand results for future use.

A high frequency energy source 30 is coupled to and controlled by theCPU 14. The source 30 is coupled to a powwer splitter 32, which dividesenergy into a plurality of lines each having the same phase and power.The phase of the energy in each line can be individually adjusted in aline stretcher or phase shifter 34. The output of each line stretcher 34is coupled to one individual applicator 36. The actual delivery of powerto the applicators 36 is controlled by switches 37 located between theline stretchers 34 and the applicators 36. The switches 37 may be simpleon-off switches, such as relays or solid state switches, or they may becontinuously variable attenuators. Like the source 30 and the phaseshifters 34, the switches 37 are preferably controlled by the CPU duringsystem operation, although the switches 37 may be manually operated.FIG. 1 shows only two line stretchers 34, applicators 36 and switches37, but an actual system 10 may employ at least four of each to providesteering in several directions.

Referring to FIG. 2, eight individual applicators 36 are shown coupledtogether in an octagonal arrangement and surrounding a circular target12. Each applicator 36 is diagrammatically represented by a rectangle.In reality, each applicator 36 would have a shape suitable for theemission of microwave EMR, several embodiments of which are shown inlater drawings. FIG. 2 is a two-dimensional representation of athree-dimensional phenomena, with both the radiators 36 and target 12extending for some distance perpendicular to the plane of the drawing.The radiation emitted from each applicator 36 is aligned so that theelectric field component is perpendicular to the plane of the drawingand the magnetic field component lies in the plane of the drawing. Thestylized wave fronts shown in FIG. 2 approximate the geometry of themagnetic field component of EMR emitted by the various applicators.

As the radiation emitted by the various applicators 36 converges on thetarget 12, it is seen that the electric fields of the radiation arelined up so that the target 12 sees, approximately, a convergingcircular wave front. The energy of the various wave fronts converges inthe center of the target 12, where the electric field addsconstructively and heats the center regions 38 of the target 12 to agreater degree than that caused by any one of the applicators 36 alone.This improved deep internal heating is caused without dangerouslyincreasing the radiant energy density at the surface of the target 12,as the incoming energy is normally spread equally over the entire targetsurface. Thus, the energy imparted to the target 12 is concentrated nearthe center, where it is desired, and minimized to the extent possible atthe target surface.

As described in connection with FIG. 1, the energy radiated by eachapplicator 36 has a constant phase relationship with that emitted by theother applicators 36. This creates a synergistic result in the centerarea 38 of the target 12, whereby the target 12 is heated to a degreegreater than that of a simple sum of the energy of the variousapplicators 36. The synergistic result will be described in more detailwith relation to FIGS. 3 and 4. With all of the applicators 36 operatingprecisely in phase, the central heating area 38 will be symmetricalaround the center point of a homogeneous target 12. If the shape orlocation of the central heating region 38 is derived to be other thansymmetrical about the center, changing the relative phase of the EMRemitted by the various applicators slightly will cause the centralheating region 38 to move generally toward the applicators 36 which arephase-lagging the remainder. By controlling the phase of energy emittedby the applicators 36 as described in connection with FIG. 1, it istherefore possible to manipulate the location of the central heatingregion 38 to best achieve the desired result. Manipulation of thecentral heating region 38 can also be accomplished through control ofthe switches 37. The power to each applicator 36 can be controlled asdesired, through either on-off switches or continuously variableswitches or attenuators as described above. Lowering or cutting offpower to individual applicators 36 changes the shape of the centralheating region 38, and the power absorbed at various points in thetarget 12.

FIGS. 3 and 4 show the mechanism by which the greatly increased powerdeposition in the central heating region 38 occurs. Considering any pairof diametrically opposed applicators 36 of FIG. 2, and a non-lossyhomogeneous target 12, the drawing of FIG. 3 shows the standing waveamplitudes of the E-field component of the EMR generated by suchopposing pair on a line through the center of the target 12. Thehorizontal axis represents the distance between the opposing applicatoremission faces, shown as points F₁ and F₂, and the vertical axisrepresents the amplitude of the alternating E-field standing wave ateach distance. The points S₁ and S₂ represent the opposite surfaces ofthe target 12 with no consideration presently being made of the E-fieldexternal to the target 12.

Because the two oncoming wave fronts are of identical frequency and havetheir E-fields aligned parallel to the center axis of the target 12, theelectic field at each point between the applicators is the sum of theE-field vectors of each wave. When the frequency of the emittedradiation is chosen so that the wavelength in the target 12 isapproximately three-fourths the diameter of the target 12, the amplitudeof the standing wave caused by two applicators 36 in the target 12 isshown in FIG. 3. The maximum amplitude is located in the center region38, with minimums being located one-fourth wavelength to either side ofthe center. The amplitude at the center is the sum of the amplitude fromeach applicator 36, which for the two opposed apertures is twice theE-field created by a single applicator 36. When more than twoapplicators 36 are used, as shown in FIG. 2 for example, the resultantE-field sum is of course larger.

Testing has shown that best results are normally obtained when thewavelength of emitted EMR is between approximately 3/4 and twice thetarget 12 diameter. This gives a relatively well-defined central heatingregion 38 and a good impedance match between the applicators 36 asdescribed below and the target 12. Thus, for a target 12 diameter d, thepreferred range of wavelengths can be found from the expression:

    0.5λ.sub.m ≦d≦1.3λ.sub.m       (1)

where _(m) is the wavelength in the tissue medium being heated. For highwater content tissues, such as muscle and blood, the wavelengths at 100,300 and 915 MHz are approximately 27, 11.9 and 4.5 cm respectively. Forlow water content tissues, the respective wavelengths at the frequenciesare approximately 106,41 and 13.7 cm. If both types of tissue arepresent in a target, it is preferable to select a wavelength whichsatisfies equation (1) for the most prevalent tissue type (normallymuscle tissue). A wavelength larger than suggested by equation 1 can beused if adequate impedance matching is obtained or provided by externalmatching techniques and nearly uniform surface vs. central heating canbe expected.

FIG. 4 shows the relative power density at each point in the target 12corresponding to FIG. 3. The power density is proportional to the squareof the electric field strength, so that the power density curve shows arelatively sharp peak in the central heating region 38 for anon-attenuating medium. Heating at any point is due to the powerabsorbed at that point, which is in turn directly proportional to thepower density at that point. Therefore, a heating cross section of thetarget has the same distribution as the power density curve of FIG. 4when heat transfer effects are neglected. However, a medium capable ofabsorbing the radiant power is attenuating and will substantially reducethe central power density peak as represented in FIGS. 3 and 4 andincrease the power density somewhat at the surface, resulting in nearuniform heating being dependent on frequency, tissue diameter, andtissue conductivity.

Since the power density is proportional to the square of the E-field, asimple additive increase in the electric field at a given point resultsin an increase in the power density at that point by the square of theelectric field. For example, in FIGS. 3 and 4 the electric field in thecentral heating region 38 resulting from 2 apertures is twice that dueto a single applicator 36. Therefore, the power density of the centralregion is 2² =4 times the power density that would be caused by a singleapplicator 36. When, as in FIG. 2, more applicators 36 are used, theincrease in power density becomes much greater than that caused by asingle applicator 36. When eight applicators 36 are used, the E-field atthe center is 8 times that caused by a single applicator 36 alone, andthe power density at the center is therefore 8² =64 times the powerdensity caused by a single applicator 36. This enormous increase inpower density, and thus power absorbed, in the center of the target 12is obtained without significantly increasing the power density at anyone point on the surface of the target 12. This phenomena, a synergisticresult due to all applicators 36 operating at the identical frequencyand with a predetermined phase relationship, allows deep heating of thetarget 12 without undesired excessive heating of the surface portions.

The above discussion of FIGS. 3 and 4 applies to non-lossy targets 12.In such targets 12 there is little energy absorption by the medium, sothat the amplitude of the EMR from any given applicator 36 isundiminished as the radiation passes through. However, actual targets 12are lossy, so that the amplitude of EMR decreases as it passed throughthe target 12. In a typical case, the amplitude of the E-field at thecenter region 38 may be approximately 1/7 the E-field amplitude at thesurface of the target 12. For the case of eight applicators 36, thepower density in the center region 38 is approximately 8² /7² orapproximately 1.3 times that at the surface of the target 12. It isimportant to note, however, that the power density in the center region12 is still sixty-four times that which would be caused by a singleapplicator 36 alone. The general shape of the E-field and power densitywave forms of FIGS. 3 and 4 still apply, with the actual peak value inthe center region 38 being diminished with a lossy medium to nearly thesame level as the surface.

The above-described synergistic increase in power density in the centralheating region 38 occurs only when all applicators 36 radiate at thesame frequency with the E-fields of the emitted EMR aligned. ThisE-field in phase alignment preferably occurs along the central axis ofthe target 12, which is perpendicular to the page as shown in FIG. 2.The E-field alignment of the various applicators 36 is preferably madeas accurate as possible. However, some misalignment can be toleratedwithout unreasonably reducing the performance of the system 10. Thevector sum of the E-field at any one point is equal to the sum of theindividual E-field vectors. If a particular applicator 36 is misaligned,the EMR power emitted by that applicator contributes less to thesynergistic power increase than an aligned field by a factor equal tothe square of the cosine of half the angle between the misalignedE-field and the remaining E-fields. For small angles the cosine is closeto one, so that small mismatches in the E-field alignment do notsubstantially degrade the synergistic power density enhancement whichoccurs in the central heating region 38.

When the various applicators 36 radiate energy at slightly differentfrequencies, it will become apparent to those skilled in the art thatthe various E-fields will not always add constructively, and the powerdensity enhancement described above will not occur. In fact, in such acase, the power density in the central region 38 will be, at best, thesimple sum of the individual applicator power densities. It is thereforeimportant that the frequency of radiation emitted by all applicators 36be absolutely identical. For this reason, the preferred embodimentutilizes a single power source 30 and a power splitter 32 whereby thepower supplied to each applicator 36 has the same frequency. While it ispossible to use multiple sources providing that they can be preciselyphase locked to emit identical frequencies, the practical considerationsto accomplish this would complicate the system and add expense withlittle benefit, thus the preferred configuration requires only a singlesource 30 and splitter 32. It is not necessary that the frequencysupplied to the applicators 36 be invariant with respect to time. Infact, it is desirable that the source 30 have a controllable frequencyso that it may be adjusted to optimize performance with various target12 materials as described above.

The shape and location of the central heating region 38 is determined bythe distribution of applicators 36 in operation, the relative phasebetween them, the diameter of the tissue, the positioning of the tissueand the EMR frequency. It has been determined that the use of four ormore uniformly spaced radiating applicators 36 will provide anapproximately circular (ellipsoidal in three dimensions) central heatingregion 38. Eight radiating applicators 36 are used in FIG. 2 instead offour, because it has been determined that the power density, and thusheating, at the surface of the target 12 is more uniform than when onlyfour radiating applicators 36 are used. An increase in the number ofapplicators 36 above eight does not appear to make a material differencein the operation of the system 10.

When all of the applicators 36 are precisely in phase, a symmetricalheating region 38 is formed in the precise center of a homogeneoustarget 12. Varying the relative phase of energy emitted by the variousapplicators 36 will cause the central heating region 38 to be shiftedsomewhat away from center toward the applicators 36 with lagging phase.The ability to alter the relative phase and amplitude between theapplicators 36 is extremely useful, for example when a non-homogeneoustarget 12 (such as an animal torso) is used. The wavelength of the EMRwill vary slightly in the different tissues of the target 12, and analteration in emitted phase can compensate for the phase shifts therebyinduced. Thus, when the target 12 has a known cross section of differenttissues having known properties, the phase between the variousapplicators 36 can be adjusted to position the central heating region 38at the desired location. Off-set positioning of the tissue can alsocompensate for non-homogeneous effects on the heat pattern.

One applicator array 40 suitable for use with the present invention isshown in FIGS. 5 and 6. This annular array 40 comprises a battery ofsixteen horn-type parallel plate waveguide antennae forming a foldeddipole array being coupled together into two layers of eight radiators36 each. For simplicity in construction, each input to the applicatorarray 40 feeds a two by two array of individual applicators 36. Thus,only four power inputs are needed for this sixteen applicator array 40.

It has been determined that the stacking of two or more applicators 36in the E-field direction (perpendicular to the page in FIG. 2) stillprovides a substantially uniform electric field vertically but reducesthe applicator size. It has also been determined that stacking of theindividual applicators 36 in an annular array 40 along the H-field (inthe plane of the page in FIG. 2) provides a substantially uniformelectric field around a target 12. Use of an array of individualapplicators 36 allows each one to be sized and constructed so as toprovide a good impedance match between the applicators 36 and the target12. A detailed discussion of methods for fabricating horn-typeapplicators 36 suitable for use with the present invention is containedin pending U.S. patent application Ser. No. 136,506, filed on Apr. 2,1980, and titled Annular Electromagnetic Radiation Applicator ForBiological Tissue And Method, which disclosure is herein incorporated asif set forth verbatim.

The applicator array 40 is surrounded by a casing 42, which serves tosupport the individual applicators 36 in place and to decrease strayradiation, which can be hazardous. As shown in FIG. 5, the casing 42 ispartially cut away, exposing portions of four separate applicators 36. Acoaxial power input line 44 is coupled to a parallel plate waveguide 46.The waveguide 46 is coupled to four feed guides 48. The feed guides 48are in turn coupled to four individual applicators 36, and have the samedimensions so that power is split evenly to the applicators 36. Each setof four applicators 36 therefore radiates energy having the same phase,power and E-field alignment.

FIG. 6 shows a top view of the applicator array 40. A target 12 issuspended interiorly of the array 40, and is surrounded by a bolus 50.The bolus 50 preferably contains deionized water, and is made from aflexible material so as to seal tightly around the target 12. Air gaps52 may be left between portions of the bolus 50 and applicators 36, orthe bolus may be manipulated to fill these gaps as desired.

The use of a bolus 50 has several important advantages. The fluidtherein can be circulated through an external heat exchanger (not shown)to cool surface regions of the target 12. When deionized water is usedin the bolus 50, there is very little power loss in the bolus 50, sothat the full power radiated by the applicators 36 is delivered to thetarget 12.

Use of the bolus 50 improves the impedance match between the applicators36 and the target 12. At the frequencies of interest, the impedance of atypial biological target 12 is approximately 44 ohms. The impedance ofthe applicators 36 and other electrical portions of the system ispreferably 50 ohms in order to be compatible with standard components.The impedance of deionized water at the frequencies of interest is alsoapproximately 44 ohms, so that all parts of the system 10 are inherentlyclosely matched. If the water-filled bolus 50 were not present, a largemismatch would occur at the radiating face of the applicators 36 and atthe surface of the target 12. This mismatch occurs because the impedanceof air is approximately that of free space, or 377 ohms. Any impedancemismatches cause reflections at the boundaries, lowering the percentageof radiated energy delivered to the target 12 and increasing strayradiation hazards.

As can be seen in FIG. 6, the applicator array 40 emits energy in thepattern discussed hereinbefore in connection with FIGS. 2 and 3,resulting in the power density pattern shown in FIG. 4. Thus, theapplicator array 40 provides heating in the central region 38 withoutdamage to the surface regions.

An alternate embodiment of an applicator suitable for use with thepresent system 10 is shown in FIG. 7. This applicator 54 is essentiallya dipole antenna pair sized for use with EMR of the frequenciescontemplated. Each arm 56 of the upper and lower radiating portions58,60 acts as a single radiator in a manner similar to that of theannular array 40 of FIGS. 5 and 6. A coaxial feed line 61 is coupled tothe center of the upper and lower radiating portions 58,60, a balun canalso be used to transform the coax-line to a balanced feed. When thisapplicator 54 is driven in a conventional manner, the E-field of theemitted radiation is aligned with the length of the arms 56.

The shape and size of the antenna arms 56 determine the optimumfrequencies of operation and impedance characteristics of the dipole 54.It has been determined experimentally that a dipole 54 having taperedarms, wherein the ratio of arm width (W) to length (L) is maintainedconstant at approximately 0.087, gives a good impedance match withremainder of the system 10. When the dipole applicators 54 are combinedinto a cylindrical array 66 as shown in FIG. 8, and a water bolus (notshown) as similar to that discussed with FIG. 6 is used, a reasonablygood 50 ohm impedance match is achieved. This approach tends to be morenarrow band than the annular array of FIG. 5, but is much simpler tobuild.

Referring to FIG. 8, four dipole pair radiators 54 are assembled on arigid, non-conducting frame 64 to form a cylindrical array 66. Eachdipole radiator 54 is separately connected to the source 30 and powersplitter 32 through a separate coaxial feed line 61. Each applicator 54launches microwave EMR towards the center of the array 66 where thetarget specimen (not shown) is located. Preferably, a de-ionized waterbolus (not shown) surrounds the target so as to better couple energyfrom the applicators 54 to the target, and to minimize reflections. Thephase of the energy to each dipole applicator 54 can be controlled tovary the location of the central heating region 38 or to compensate forwavelength variations in non-homogeneous targets as described inconnection with FIG. 2.

Referring to the alternate embodiment of FIG. 9, a cylindrical dipoleapplicator 68 comprises two coaxial conducting cylinders 70 placed closetogether. These concentric cylinders 70 act as a single dipoleapplicator wherein the radiating arms comprise a flat radiating sheetwhich has been folded around to make contact with itself. Thecylindrical dipole 68 will radiate toward its central axis, andconstructive interference of the E-field will cause the synergisticallyincreased power absorption in the central region 38 as described withthe previous applicators. A single coaxial feedline 72 is sufficient todrive the cylindrical dipole 68. However, there will be some phase lagin the EMR emitted from the portions of the dipole 68 diametricallyopposite the feedline 72. This will cause the central heating region 38to shift somewhat away from the feedline 72 contacts. While this may bedesirable in some cases, the preferred embodiment includes four coaxialfeedlines 72 equally spaced around the dipole 68. When all fourfeedlines 72 are driven at the same phase, the central heating region 38will be centered around the axis of the cylindrical dipole 68. Somemanipulation of the central heating region 38 location can be made byvarying the phase to the coaxial feed lines 72, but in general thedegree of control will be less than that experienced with either thedipole array 66 or the horn radiator array 40.

Since the effective radiating aperture width of the cylindrical dipole68 is equal to its circumference, and the height is limited by the sizeof the target to typically two feet or less, inherent impedance matchingas was obtained with the dipole array 66 is difficult to achieve. Sincethe cylindrical dipole radiator 68 will not be inherently matched withthe impedance of the remainder of the system at most desirablefrequencies, a conventional impedance matching device (not shown) shouldbe used to minimize losses and reduce reflected power.

Both the folded dipole array 62 and the cylindrical dipole 68 emitradiation from both their inner and outer surfaces. An internal waterbolus will increase the proportion of radiation emitted centrally, dueto the better impedance match. In order to reduce further the hazard ofstray radiation, an outer conducting cylinder (not shown) can be placedaround the cylindrical dipole 68 or dipole array 66. This outer shieldcan be grounded or left floating in order to reflect such radiation andreduce the outwardly emitted radiation. The reflecting shield must bespaced a sufficient distance from the cylindrical dipole 68 or dipolearray 66 so that the ground plate will not capacitively load theaperture so much as to interfere with the primary emitted radiationdistribution and reduce heating in the central region 38 or causeundesired non-central heating. In the preferred embodiment, the outerconducting cylinder is grounded and displaced so that minimal powerpattern changes occur. Normally the space between the outer cylinder andthe dipoles is filled with air or another low dielectric materialthereby reducing the amount of energy coupled to the shorting cylinder.The grounding of the outer conducting cylinder is preferably done with asecond coaxial outer shield dielectrically spaced from and outside ofthe coaxial outer conductor connecting to the dipole radiators.

In order to provide an effective hyperthermia treatment, the operatormust be able to accurately determine the internal status of the target12. For a living target 12, monitoring the vital signs gives a generalindication of the health of the target 12 and indicates adverse eventsaffecting its health. However, these signs, such as pulse, respiration,blood pressure and oral temperature, do not indicate whether enough heatis being applied to the region of interest to be effective.

Two additional measurements provide a fairly complete picture of theinternal local effects of the hyperthermia treatment. The first of theseis the measurement of actual temperature at selected points within thetarget 12. A real time thermal profile allows the operator to determinewhether the desired regions of the target 12 are being heated totemperatures which are medically effective. Such a profile also allowsthe operator to ensure that no unwanted heating occurs in undesiredportions of the target 12.

Referring to FIG. 10, a means for generating a thermal map of thetarget's interior is shown diagrammatically. Two individual probes 22are shown, although four to eight are commonly used when the system 10is in operation. The probes 22 are invasively placed in the target 12.Catheters 76 are initially inserted into the target 12 in apredetermined location. The catheters 76 have closed tips so that nofluids from the target 12 flow into the catheter 76. The catheters 76have surface indicating guides 78 which are pressed flush with thesurface of the target 12 to ensure proper positioning. Coupled to thecatheter 76 at the surface guide 78 is a flexible but fairly stiffhollow tube 80. This hollow tube or casing 80 is firmly coupled to asolid mounting plate 82.

A thermal sensor 84 is located inside the catheter 76. Such sensor 84 ispreferably a non-metallic high resistance sensor, and should be of atype which will not affect or be affected by incident microwaveradiation. The sensor 84 must reflect the true temperature in the regionof its placement and should be free of distortion from the microwaveradiation which falls upon it. Sensor leads 86 have a high electricalresistance so as to be unaffected by the electromagnetic field, and arewrapped in a dielectric material so as to form a single stiff butsomewhat flexible lead 86. In the alternative, an optical thermal sensor84 using fiber optic leads 86 may be used, as it is unaffected bymicrowave radiation. The leads 86 run through the catheter 76 and thecasing 80, and pass through the fixed plate 82. The interior diameter ofthe casing 80 is slightly larger than the diameter of the stiffened lead86 so as to allow longitudinal movement without binding.

Since the thermal sensor leads 86 are stiff, movement of such leads 86will vary the position of the sensor 84 within the catheter 76. Thisallows temperature measurements to be made at several positions alongthe catheter 76 during operation of the system 10, thus maximizingthermal gradient information while minimizing the number of invasivedevices required.

The probe leads 86 pass through the plate 82, and are coupled togetherby a collar 88. Two capstan rollers 90 keep the leads firmly in contactwith a sheave 92, which in turn is driven by a stepping motor (notshown) or other means which allows the position of the leads 86 to beprecisely determined. This precisely determines the position of thevarious temperature sensors 84. As shown in FIG. 10, four measurementpositions P for the sensors 84 correspond with four predetermined sheave92 positions.

During operation of the system 10, the CPU 14 is programmed to recorddata, taken by the thermal sensors 84, in the memory 28. The position ofthe stepping motor is controlled by the CPU 14, and the sensors 84 maybe positioned in a predetermined measurement sequence, or a specialsequence if the operator determines that an anomaly appears to bepresent and desires specific readings pertaining thereto.

The central processor 14 causes the thermal information taken duringoperation of the system 10 to be stored in long-term memory (e.g., disk)for later analysis. The system 10 also uses this information in afeedback loop to control the operation of the power source 30 andsplitter 32. If heating should rise to dangerous levels, the powerapplied to the target 12 can be decreased. Likewise, if the heating isinsufficient, extra power can be applied. The CPU 14 can also display atemperature vs position plot of several of these scanning sensors 84 forthe operators to view.

In addition to the real time thermal condition of the target 12,information describing the real time power absorption patterns by thevarious tissues can be supplied by the present system 10. The powerabsorbed by a given region of tissue is directly proportional to thepower density in that region. As discussed in connection with FIGS. 3and 4, the power density is proportional to the local squared electricfield of the microwave radiation emitted by the applications 36. In anon-homogeneous target 12 such as a biological specimen, the powerdensity in a particular tissue is also proportional to the conductivityof that tissue and the local squared E-field.

When the system 10 is used to generate hyperthermia in a human torso, itis preferable to take a computerized axial tomographic (CAT) scan of theregions to be heated. The results of the CAT scan are reduced to a twoor three dimensional model suitable for computer processing. The modelconsists of a relatively large number of small "cells", or regions. Forexample, 20-50 cells may be used to model a human heart.

Each cell consists of only one tissue type, and is assigned the knownproperties ascribed to that tissue type. For example, at frequencies inthe approximate range of 50-900 MHz, high water content tissues such asmuscle and blood have a real part of permittivity approximately 50-90times that of air, and a conductivity in mhos/meter in the region of0.6-1.5. Precise values depend on the precise tissue type. Deionizedwater has a real permittivity of approximately 80 times that of air anda conductivity of zero. The close match in permittivity causes goodimpedance matching with high water content tissues and the zeroconductivity ensures no power loss by EMR passing through, which areamong the reasons a water bolus 50 surrounding the target 12 ispreferred as described above.

Low water content tissue such as fat and bone have real permittivitiesand conductivities much lower than high water content tissues, so thatpower absorbed by these tissues is much lower for the same appliedE-field than for high water content tissues. Use of both the real partof permittivity and conductivity is necessary to model the target 12.Experience has shown that by initially assuming that the target 12 isrepresented by an equivalent homogeneous media the calculations can besimplified, giving a good approximation of the E-field distributions inthe target 12. The conductivity of each cell is then used to predict theestimated power density by Ohm's law. Further refinement of the localE-fields can be estimated by quasi-static approximations using thevarious permittivities.

Given a model of the target 12 to be heated, and the properties of thevarious tissues encountered, predictions are made of the powerabsorption patterns expected with the application of EMR. The predictionfunctions assume that each cell is homogeneous, and use conventionaliterative methods to arrive at a solution. Solutions are calculated fora range of frequencies, amplitude, and phase relationships between theapplicators 36. For each operating situation, first the E-field solutionis generated, followed by the corresponding power density solution. Thesolution templates are stored in the system memory 28 for reference andcomparison by the CPU 14 during operation of the system 10.

It is theoretically possible to have the CPU 14 generate the appropriatesolutions during system 10 operation. However, solving the model forvarious operating conditions consumes enormous amounts of processor timegiven the present state of the art, which would not allow the system 10to operate in real time. By generating solution templates in advance,the CPU 14 can quickly scan for the predicted solution that most closelymatches known conditions. The frequency, amplitude, and phase of emittedradiation can be controlled as necessary in order to provide the desiredpower density patterns.

If the heat transfer characteristics of the targert 12 are known,temperature distribution templates can also be generated in advance. Dueto the complexity of blood flow and heat transfer in biological systems,it is not presently possible to accurately predict temperaturedistribution patterns in living biological targets 12 generally.However, after several hyperthermia sessions have been undergone with aparticular target or class of targets, the stored results allow betterestimates of heat transfer patterns. Also, the system 10 monitors thereal time temperature distributions, and is able to compensate forunforeseen thermodynamic imbalances.

As described above, the power density in a region is proportional to thelocal E-field squared and the tissue conductivity. Since the spatialdistribution of tissue types is known from the CAT scan, measurement ofthe E-fields in various locations determines the power density. For agiven target 12, a given E-field distribution external of the target 12under known operating conditions (frequency, phase, specimen location,and so forth) implies a unique internal E-field distribution. Bymeasuring the E-field distribution near the surface of the target 12,and matching these values with the closest prediction template, theE-fields within the target 12 can be determined without the necessityfor invasive measuring devices.

Referring to FIG. 11, a cross-section of a human torso 94 is shownlocated within an annular array 40 of the type described in FIGS. 5 and6. The simplified cross-section of the torso 94 consists of a chest wall96, ribs 98, lung tissue 100, spine 102, heart 104 and a major artery106. A bolus 50, preferably containing deionized water, surrounds thetorso 94 and fills the spaces between it and the radiating surfaces 108.Located adjacent to the torso 94 are a plurality of E-field detectors110. These detectors 110 are small semiconductor diode detectors whichdo not distort the electric field which they are measuring. Since theE-field of the radiation emitted by all of the radiating surfaces 108 isparallel, and perpendicular to the page of the drawing, the E-fielddetectors 110 are also aligned perpendicular to the page of the drawing.For simplicity and accuracy in placement of the detectors 110, they maybe coupled to the bolus 50 adjacent to the torso 94. The externalE-field amplitudes as measured by these detectors 110 allowdetermination of the internal E-field and power density distributions asdescribed above.

FIG. 12 is a schematic diagram of the electric field detector 110. InFIG. 12 the electric field detectors or electric field probe arrays aremade up of resistive conductors 124 that are connected between thecathode and anode of diodes 125 as shown. The diodes are collectivelyconnected by cable 126 which is in turn connected to cable 123 at somedistance. Cable 123 includes a resistance 122 in series and a shuntedvariable resistance 121 across the two terminals 120 which terminate thecable 123.

These electric field probes include the resistive dipole which, in thepreferred embodiment, is larger than the size of any superficialpertubating nonhomogeneous internal structure (typically 8 cm long), butsmaller than one third of a wavelength of any dominant tissue. In anormal human body the muscle tissue is the dominant tissue. Thesecharacteristics assure that at the frequencies applied to the regions,the dipoles are too short to resonate and thus reduce any effects offield and frequency dependent detection. This dipole resistance alsoreduces the probe effect on the local field and reduces excessive probeheating. Such a probe will measure the average field intensity eventhough several field intensity variations may be present across thedipole region. An even better comparison is made when several probes areplaced at various locations within the heating field, assuming that thetissue variations are similar in each site where the probes are placed.However, where underlying structures are fairly large (such as an airfilled stomach), these structures could cause modifications in theexternal field. To counteract this, multiple probes placed in line andalong the dominant E-field direction sample a wider tissue region toallow field variations to be averaged. This same effect can also beaccomplished by connecting these multiple probes in series so thedetected voltage of each adds to the others in line with it. Such anarray would provide a single measurement of the relative field over alarge tissue region within the primary heating area. This level can thenbe easily compared with the level of other arrays similarly placed atother positions along the external body surface within the heatingfield. However, it should be obvious to one skilled in the art thatthese probes may be placed internal to the body within cathetersinserted in natural openings or surgical openings.

In the preferred embodiment the probes were intended for use atfrequencies below 110 megahertz for hyperthermia treatments of a humantrunk however smaller body regions such as limbs may require higherfrequencies. The muscle wavelength at 110 megahertz is approximately 30centimeters. The diodes used in the probes are standard signal ordetecting diodes which provide good result in estimating the overallaverage field level. In the preferred embodiment, the diodes used are tobe operated in their linear region or zero bias Schotky diodes should beused in order to provide an output voltage that will represent the sumof the square of the electric fields of each of the detecting dipoles.If the diodes operate in a nonlinear region, the detected voltage levelof a dipole detecting a high power will not affect the final voltageoutput as much as arrays operating in the linear region and the resultwill be slightly inaccurate. Therefore, operation in the linear regionis preferred to maintain accuracy.

Also in this preferred embodiment, the probe array includes multiplediode loaded dipoles in a linear array. Separate leads can connect toeach diode for a separate read out of each diode. Alternatively, thediodes may be placed in series in order to sum the voltages which wouldrepresent the average or they may be placed in parallel in order to sumthe currents to represent the average. This complete assembly is thenplaced inside a dielectric heat shrink tube with the tip sealed. Theterminal end in the preferred embodiment contains a high impedance (over30 megaohms) voltage divider with a potentiometer to calibrate thevoltage range between zero to five volts when the probe array is placedin a known field.

The dipole can be constructed in either of two forms: a single short (9millimeter total length) silicon diode, or this same diode attached inline to a 7.6 centimeter carbon-loaded dipole having a dipole resistanceof 640 ohms per centimeter of length.

Three areas of difficulty with these probes may be encountered. First,the dipole output voltage may be too large. The second is thatvariations may exist between probes requiring individual calibration.The third is that the very high input impedance of the monitor combinedwith the large lead capacitance may result in very slow response time.In order to solve these problems associated with a long dipole, aresistive attenuator may be installed at the connector (such asResistances 121 and 122 in FIG. 12). In the preferred embodiment, thisattenuator includes two 15 megaohm resistors (shown as Resistance 122)in series with the leads and a 5 megaohm potentiometer 121 shuntedacross the two 15 megaohm resistors. The potentiometer 121 may beadjusted to provide the proper voltage output when the probe is placedin a known electric field. The reduced detection voltage together withthe resistive loading will solve the high output problem, improve theresponse time and provide calibration of each individual probe.

The probe lead in the preferred embodiment is 1.8 meters long and 1.3millimeters wide and 0.3 millimeters thick, and is enclosed in a heatshrink plastic tube 2 millimeters in diameter. The two conductors inthis lead are 1.3 millimeter wide carbon-impregnated Teflon ribbonshaving a resistance of 6800 ohms per centimeter of length. These tworibbons are separated by a layer of insulating materials 0.1 millimetersthick and having a dielectric of 3.0. The capacitance of 1 millimeterlength of this probe lead was determined to be 345 picofarads. Theparallel capacitance between leads and the lead resistance provides avery uniformly distributed filter for suppressing common modeinterference.

FIG. 13 illustrates a electric field detection system including at leastone electric field probe 200 as described above connected to a receivingapparatus 201 that detects a voltage produced by the probe. Thisillustrates the general application of these probes connected to areceiving apparatus for detecting the presence of electric fields.Several probes may be connected to detect the presence of electricfields at different physical locations.

Although a preferred embodiment has been described in detail, it shouldbe understood that various substitutions, alternations, andmodifications may become apparent to those skilled in the art. Thesechanges may be made without departing from the spirit and scope of theinvention as defined by the appended claims.

What is claimed is:
 1. A system for heating a target withelectromagnetic radiation, comprising:a central control unit; a sourcecoupled to said control unit, said source generating electromagneticenergy at a frequency and power in response to said control unit; meanscoupled to said source for transferring energy from said source to saidtarget; and input means coupled to said control unit for indicating thestatus of the target, wherein operation of the system is controlled bysaid control unit as a function of the target status, said input meansincluding a plurality of electric field probe arrays, each electricfield probe array including a plurality of dipole probes aligned along acommon axis, each including a diode connected between two linearconductors and said diodes of each probe array collectively connected inseries.
 2. A system according to claim 1, wherein said dipole probes arecollectively connected to at least two resistive leads having aresistance per unit length value greater than that of one of said dipoleprobe conductors.
 3. A system according to claim 2, wherein saidconductors are of a size smaller than one third of a wavelength of astructure contained within said target.
 4. A system according to claim3, wherein said electric field probe arrays include catheters forinsertion into the target.
 5. A system according to claim 4, whereinsaid source frequency is below 110 megahertz.
 6. A system according toclaim 5, wherein said diodes of each of the electric field probe arraysoperate in a linear region of their characteristic curves.
 7. A systemaccording to claim 6, wherein each diodes include zero biased Schotkydiodes.
 8. A system according to claim 2, wherein the at least tworesistive leads have a resistance characteristic of at least 600 ohmsper centimeter of length.
 9. A system according to claim 8, wherein eachof said arrays is sealed with a dielectric insulator.
 10. A systemaccording to claim 9, wherein each of said electric field probe arraysincludes a calibration means for adjusting the voltage output of thearray.
 11. An electric field probe array comprising:a plurality ofdipole probes aligned along a common axis, each including a diode andtwo linear conductors, said diode connected between the two linearconductors, said diodes of each probe of the array collectivelyconnected in a series with two resistive leads connected to opposingends of said series connection, said resistive leads having a resistanceper unit length of a value greater than that of one of the dipoleconductors.
 12. An electric field probe array according to claim 11,wherein said two linear conductors are each of a size smaller than 1/3of a wavelength of a structure contained within a target.